Gamma ray detectors may be used in different applications, such as in positron emission tomography (PET) systems. PET systems perform nuclear medicine imaging that generates a three-dimensional image of functional processes within a body. For example, a PET system generates images that represent the distribution of positron-emitting nuclides within the body of a patient. When a positron interacts with an electron by annihilation, the entire mass of the positron-electron pair is converted into two 511 keV (i.e., annihilation) photons. The photons are emitted in opposite directions along a line of response. The annihilation photons can be detected by detectors that are placed along the line of response on a detector ring. As shown in FIG. 1 (PRIOR ART), a PET detector 10 includes a plurality of modules 22 that are arranged in a ring. Each detector module 22 is assembled from a plurality of detector units, or blocks 24. When the annihilation photons arrive and are detected at the detector blocks at the same time, this is referred to as coincidence. An image is then generated based on the acquired photon detection data that includes the annihilation photon detection information.
Recently, the silicon photomultiplier (SiPM) has been widely used in PET. The SiPMs are tiled arrays of up to tens of thousands of avalanche photodiodes of typical size between about 10 microns to about 100 microns, connected in parallel on a common silicon substrate and working on common load. The output of an SiPM device is typically connected to a buffer amplifier, which can be implemented as a transimpedance amplifier. Compared to a conventional vacuum photomultiplier tube (PMT), the SiPM has the advantages of compact size and allowing mass production with reduced variability and lower cost per unit of photosensitive area. Other superior characteristics are related to operation and performance, such as higher photon detection efficiency (PDE), lower bias voltage, better timing resolution, and insensitivity to magnetic fields. However, the SiPM has the disadvantages of higher dark count rate, slower fall time of output pulse relative to the PMT, and signal-correlated spurious effects such as cross-talk and after-pulsing. These effects are cumulative with the number of SiPM devices connected into a PET detector block, and result in significant timing resolution degradation as the block size is increased.
In the detector block, the crystal is optically coupled to the SiPMs. For the detector design with the scintillator crystal directly coupling to the SiPM device and 1-to-1 readout, the minimal loss and propagation of optical photons between the scintillator and the SiPM, and negligible crosstalk among SiPM devices results in better timing performance. The detector, however, requires numerous readout channels, as well as consumes a significant amount of power to incorporate complete readout electronics for each SiPM. It is thus very complex, difficult, and expensive to implement.
In addition, a fraction of the incident 511 keV annihilation photons produce multiple interactions in multiple crystals and thus readout channels of the corresponding SiPMs due to Compton scattering or optical photon spreading between crystals (light sharing). In a PET detector, high sensitivity is necessary for obtaining images with reasonable signal-to-noise ratio (SNR). The sensitivity of a PET scanner is primarily determined by the efficiency of the detection system (e.g. crystal thickness and the solid angle coverage). An enhanced detection system is needed that recovers Compton scattering events or light-sharing events in the scintillation block while maintaining good timing resolution. The following disclosure will address the needs as described above to provide an improved timing performance and detection efficiency. Aspects of the invention will process noisy and slow signals generated by SiPM devices, and greatly reduce the number of electronic channels to bring down the cost while offering excellent timing capabilities, e.g. sub-250 ps for time of flight positron emission tomography (TOF-PET).